Implantable stimulation devices generate and deliver electrical stimuli to nerves and tissues for the therapy of various biological disorders, such as pacemakers to treat cardiac arrhythmia, defibrillators to treat cardiac fibrillation, cochlear stimulators to treat deafness, retinal stimulators to treat blindness, muscle stimulators to produce coordinated limb movement, spinal cord stimulators to treat chronic pain, cortical and deep brain stimulators to treat motor and psychological disorders, occipital nerve stimulators to treat migraine headaches, and other neural stimulators to treat urinary incontinence, sleep apnea, shoulder subluxation, etc. The present invention may find applicability in all such applications and in other implantable medical device systems, although the description that follows will generally focus on the use of the invention in a Bion® microstimulator device system of the type disclosed in U.S. Patent Publ. No. 2010/0268309. The invention can also be used in a Spinal Cord Stimulator (SCS), such as is disclosed in U.S. Pat. No. 7,444,181, for example.
Microstimulator devices typically comprise a small, generally-cylindrical housing which carries electrodes for producing a desired stimulation current. Devices of this type are implanted proximate to the target tissue to allow the stimulation current to stimulate the target tissue to provide therapy for a wide variety of conditions and disorders. A microstimulator usually includes or carries stimulating electrodes intended to contact the patient's tissue, but may also have electrodes coupled to the body of the device via a lead or leads. A microstimulator may have two or more electrodes. Microstimulators benefit from simplicity. Because of their small size, the microstimulator can be directly implanted at a site requiring patient therapy.
FIG. 1 illustrates an exemplary implantable microstimulator 100. As shown, the microstimulator 100 includes a power source 145 such as a battery, a programmable memory 146, electrical circuitry 144, and a coil 147. These components are housed within a capsule 202, which is usually a thin, elongated cylinder, but may also be any other shape as determined by the structure of the desired target tissue, the method of implantation, the size and location of the power source 145, and/or the number and arrangement of external electrodes 142. In some embodiments, the volume of the capsule 202 is substantially equal to or less than three cubic centimeters.
The battery 145 supplies power to the various components within the microstimulator 100, such the electrical circuitry 144 and the coil 147. The battery 145 also provides power for therapeutic stimulation current sourced or sunk from the electrodes 142. The power source 145 may be a primary battery, a rechargeable battery, a capacitor, or any other suitable power source. Systems and methods for charging a rechargeable battery 145 will be described further below.
The coil 147 is configured to receive and/or emit a magnetic field that is used to communicate with, or receive power from, one or more external devices that support the implanted microstimulator 100, examples of which will be described below. Such communication and/or power transfer may be transcutaneous as is well known.
The programmable memory 146 is used at least in part for storing one or more sets of data, including electrical stimulation parameters that are safe and efficacious for a particular medical condition and/or for a particular patient. Electrical stimulation parameters control various parameters of the stimulation current applied to a target tissue including the frequency, pulse width, amplitude, burst pattern (e.g., burst on time and burst off time), duty cycle or burst repeat interval, ramp on time and ramp off time of the stimulation current, etc.
The illustrated microstimulator 100 includes electrodes 142-1 and 142-2 on the exterior of the capsule 202. The electrodes 142 may be disposed at either end of the capsule 202 as illustrated, or placed along the length of the capsule. There may also be more than two electrodes arranged in an array along the length of the capsule. One of the electrodes 142 may be designated as a stimulating electrode, with the other acting as an indifferent electrode (reference node) used to complete a stimulation circuit, producing monopolar stimulation. Or, one electrode may act as a cathode while the other acts as an anode, producing bipolar stimulation. Electrodes 142 may alternatively be located at the ends of short, flexible leads. The use of such leads permits, among other things, electrical stimulation to be directed to targeted tissue(s) a short distance from the surgical fixation of the bulk of the device 100.
The electrical circuitry 144 produces the electrical stimulation pulses that are delivered to the target nerve via the electrodes 142. The electrical circuitry 144 may include one or more microprocessors or microcontrollers configured to decode stimulation parameters from memory 146 and generate the corresponding stimulation pulses. The electrical circuitry 144 will generally also include other circuitry such as the current source circuitry, the transmission and receiver circuitry coupled to coil 147, electrode output capacitors, etc.
The external surfaces of the microstimulator 100 are preferably composed of biocompatible materials. For example, the capsule 202 may be made of glass, ceramic, metal, or any other material that provides a hermetic package that excludes water but permits passage of the magnetic fields used to transmit data and/or power. The electrodes 142 may be made of a noble or refractory metal or compound, such as platinum, iridium, tantalum, titanium, titanium nitride, niobium or alloys of any of these, to avoid corrosion or electrolysis which could damage the surrounding tissues and the device.
The microstimulator 100 may also include one or more infusion outlets 201, which facilitate the infusion of one or more drugs into the target tissue. Alternatively, catheters may be coupled to the infusion outlets 201 to deliver the drug therapy to target tissue some distance from the body of the microstimulator 100. If the microstimulator 100 is configured to provide a drug stimulation using infusion outlets 201, the microstimulator 100 may also include a pump 149 that is configured to store and dispense the one or more drugs.
Turning to FIG. 2, the microstimulator 100 is illustrated as implanted in a patient 150, and further shown are various external components that may be used to support the implanted microstimulator 100. An external controller 155 may be used to program and test the microstimulator 100 via communication link 156. Such link 156 is generally a two-way link, such that the microstimulator 100 can report its status or various other parameters to the external controller 155. Communication on link 156 occurs via magnetic inductive coupling. Thus, when data is to be sent from the external controller 155 to the microstimulator 100, a coil 158 in the external controller 155 is excited to produce a magnetic field that comprises the link 156, which magnetic field is detected at the coil 147 in the microstimulator. Likewise, when data is to be sent from the microstimulator 100 to the external controller 155, the coil 147 is excited to produce a magnetic field that comprises the link 156, which magnetic field is detected at the coil 158 in the external controller. Typically, the magnetic field is modulated, for example with Frequency Shift Keying (FSK) modulation or the like, to encode the data. For example, data telemetry via FSK can occur around a center frequency of f1=125 kHz, with a 129 kHz signal representing transmission of a logic ‘1’ and 121 kHz representing a logic ‘0’. (This frequency f1 will be discussed as a single frequency defined by its center, but this is merely for convenience; in reality, this frequency has a bandwidth as necessary for FSK communication, and should be so interpreted).
An external charger 151 provides power used to recharge the battery 145 (FIG. 1). Such power transfer occurs by energizing the coil 157 in the external charger 151, which produces a magnetic field comprising link 152, which occurs with a different frequency (f2=80 kHz) than data communications on link 156. This magnetic field 152 energizes the coil 147 through the patient 150's tissue, and which is rectified, filtered, and used to recharge the battery 145. Link 152, like link 156, can be bidirectional to allow the microstimulator 100 to report status information back to the external charger 151. For example, once the circuitry 144 in the microstimulator 100 detects that the power source 145 is fully charged, the coil 147 can signal that fact back to the external charger 151 so that charging can cease. Charging can occur at convenient intervals for the patient 150, such as every night.
FIG. 3A shows the communication and charging circuitry 101 within microstimulator 100 that is coupled to coil 147. Such circuitry is explained in detail in the '992 Publication, with which the reader is assumed familiar, and thus is only briefly explained here.
As explained in the '992 Publication, the circuitry of FIG. 3A is beneficial because it uses a single coil L1 (147) for receiving a magnetic charging field 152 from the external charger 151, and for transmitting and receiving data telemetry 156 to and from the external controller 155. (The external charger 151 and external controller 155 are shown in FIG. 3A as one integrated unit for simplicity).
Coil 147 is connected at one end through transistor switch M1 to a voltage, Vbat, provided by the battery 145 in the microstimulator 100. Coil 147 is connected at its other end through transistor switch M2 to ground. Tank capacitor C1 is connected in parallel with coil 147, and tunes the coil to a particular frequency for transmitting or receiving data telemetry to and from the external controller 155 (e.g., approximately f1=125 kHz). A series combination of a tuning capacitor C2 and transistor switch M3 are also connected in parallel to coil 147. Transistor M3 is turned on during receipt of a magnetic charging field along link 152 from the external charger 151 to tune the coil to the frequency of the magnetic charging filed (e.g., approximately f2=80 kHz). Also connected in parallel with coil 147 is a full bridge rectifier formed of diodes D1-D4 for producing DC voltage Vout. A half bridge rectifier or even a signle diode rectifier could also be used. A transistor switch M4 is also connected between the rectifier circuitry and ground.
DC voltage Vout is received at storage capacitor C3, which filters and smoothes the voltage before being passed to battery charging circuitry 92. Battery charging circuitry 92 is used to charge the battery 145 in a controlled fashion. If needed, a Zener diode D5 or other suitable voltage clamp circuit may be connected across storage capacitor C3 to prevent Vout from exceeding some predetermined value.
FIG. 3B shows the status of transistor switches M1-M4 for the energy receive, data receive, and data transmit modes. As shown, to operate in an energy receive mode, the circuit will turn switches M1, M2 and M4 OFF, and will turn switch M3 ON. Turning M3 ON includes tuning capacitor C2 in parallel with tank capacitor C1, which, in conjunction with the inductance formed by the coil 147, forms a resonant circuit which is tuned to the frequency of the magnetic charging field (f2=80 kHz). The circuit of FIG. 3A may also operate in a data transmit mode during charging by employing back telemetry known as Load Shift Keying (LSK), in which case transistor M4 is modulated with the data to be transmitted back to the external charger 151.
For the circuit of FIG. 3A to operate in a data receive mode, the circuit will turn switches M1, M3 and M4 OFF, and will turn switch M2 ON. Turning M3 off excludes capacitor tuning C2 from the resonant circuit, whose tuning is thus governed by coil 147 and tank capacitor C1. With tuning capacitor C2 excluded, the resonant circuit is tuned to a higher frequency matching the operation of the external controller 155 (f1=125 kHz). Turning M2 ON grounds the resonant circuit, which provides an input to the receiver, which demodulates the received data (DATA RCV). The receiver can either comprise a differential input as illustrated in solid lines in FIG. 3A, or can comprise a single-ended non-differential input in which one of the inputs is grounded, as shown in dotted lines in FIG. 3A.
As further shown in FIG. 3B, the circuit of FIG. 3A may also operate in a data transmit mode by turning switches M3 and M4 OFF, by modulating switch M2 with a data signal (DATA XMIT), and by turning switch M1 ON. Under these conditions, the resonant circuit is once again, by virtue of transistor M3 being OFF, tuned to the higher frequency (f1=125 kHz), and will broadcast a signal to the external controller 155 along link 156 accordingly, with the energy for the radiation being supplied from the battery voltage, Vbat, via transistor M1. The transmitter receiving the data to be transmitted (DATA XMIT), is shown coupled to transistor M2, but could also couple to transistor M1.
Thus, it is seen that by selectively controlling the state of the switches M1-M4, the circuit of FIG. 3A may operate in different modes, using only a single coil 147. Such modes may be invoked in a time-multiplexed manner, e.g., with a first mode being followed by a second mode, depending upon the particular application at hand. Control signals M1-M4, as well as DATA XMIT, are ultimately issued by a microcontroller (or, more generically, control circuitry 160) in the microstimulator 100, and DATA RCV is received by that microcontroller.
While the versatility of the single-coil, multi-function circuit of FIG. 3A is desirable, the inventors recognize drawbacks. One drawback is that storage capacitor C3 loads the resonant tank circuit (coil L1 147 and tank capacitor C1) during periods when the circuitry transmits data. As discussed earlier, during data transmission, switch M1 is closed while switch M2 is modulated with the data signal, which causes the tank circuit to resonate, thus forming an AC voltage, Vtank, with a center frequency of approximately f1=125 kHz. This alternating voltage in the tank circuit also appears across the full bridge rectifier (D1-D4). Because switch M1 is closed, the top node of the tank circuit, which node is connected to the switch M1, will remain fixed to approximately Vbat. As a result, there will be some charge leakage from this node to the storage capacitor C3 via diode D3. Because switch M2 is modulated, the voltage at the bottom node of the tank circuit, which node is connected to the switch M2, will vary between ground and Vbat. Thus, depending upon the instantaneous voltages at the bottom node and Vout, diode D4 may also become forward biased and leak charge into the storage capacitor C3. Note that diodes D1 and D2 do not conduct because they remain reversed biased. Thus, some of the charge generated in the resonant tank circuit is leaked into the storage capacitor C3, which loads the resonant tank circuit. (Other components on the DC side of the rectifier such as the battery charging circuitry 92 and the battery 145 may be disconnected or disabled during telemetry, and in any event do not appreciably load the tank circuit).
The inventors have noticed that loading of the tank circuit by the storage capacitor C3 has undesirable effects. The first relates to the speed at which the RF signal transmitted by the tank circuit—i.e., the RF signal comprising communication link 156—can reach its full strength. The strength of the RF signal is primarily governed by the magnitude of Vtank. But leakage to the storage capacitor C3 via the full bridge rectifier impedes a full strength RF signal, at least initially. This is because storage capacitor C3 is initially not charged, and such lack of charge promotes leakage through the diodes D3 and D4 as previously discussed. Eventually such leakage will charge the storage capacitor C3, which will tend to reduce the leakage through the diodes, at which point the RF signal will be at full strength. The effect is that when circuit of FIG. 3A begins to transmit data, an initial portion of the data will not be transmitted with a full strength RF signal. This makes reception of this signal at the external controller 155 more difficult to resolve, resulting in corrupted data or no data at all. Experimental results show that the length of time for the tank circuit to transmit with a full strength RF signal is approximately 2 ms. At typical data transmission rates of 4 Kbps, this delay can contribute to significant data transmission errors affecting 8 bits of information in this example. Moreover, even if storage capacitor C3 is fully charged, there can still be some leakage through the diodes in the rectifier, and hence some coupling of the storage capacitor C3 to the tank circuit, which impeded RF signal strength and detunes the tank circuit.
A second undesirable effect is that loading of the tank circuit alters its resonant frequency, especially during the time when the RF signal strength is increasing towards its maximum value. This occurs because leakage through the didoes D3 and D4 effectively places storage capacitor C3 in parallel with the tank circuit. This increases the effective capacitance of the tank circuit, which decreases its resonant frequency. In short, coupling of the storage capacitor C3 detunes the tank circuitry to less than the optimal center value of f1=125 KHz. Again, such detuning can affect the reliability of data transmission.
Another drawback of the circuit of FIG. 3A relates to switch M3 on the AC side of the rectifier. Vtank can comprise a relatively high alternating voltage, and switch M3 is therefore subject to large swings in voltage. This makes implementing and controlling switch M3 rather difficult, and can result in increased complexity, size, and cost of the circuitry.
This disclosure presents solutions to the aforementioned and other shortcomings of the prior art.